Laser therapy instruments are used, e.g., for correcting an ametropia of the human eye by a laser-surgical operation of the cornea. For this purpose, a lid is formed on the outer surface of the cornea, which lid is attached along one edge and is therefore known as a flap, and the thickness of which is substantially smaller then the thickness of the cornea. For correction, this flap is folded back, and from the surface of the region of the cornea that is now exposed, tissue is removed thereupon by application of a laser beam pulsed in the femtosecond range, in order to change the curvature of the cornea. Such an instrument is described in DE 10 2005 013 949 A1.
By contrast, a laser system described in DE 10 2008 027 358 A1 is intended for the analysis and treatment of the crystalline lens. Here, laser radiation also pulsed in the femtosecond range is focussed on to selected target spots in the region of the crystalline lens. At this wavelength, the detection of the laser light backscattered in the crystalline lens is possible at the greatest accuracy, and a refractive-surgical therapy of the crystalline lens can be performed with high precision.
Typically, in either case the interaction between the ultrashort laser pulses and the tissue takes place in a small volume, hereafter referred to as focus volume. Situated within the focus volume is the interaction zone, in which the structural change, section or removal of the tissue takes place. The laser focus must be precisely positioned at the locus of the desired interaction. This is done with an optical focusing system, which projects the parallel laser beam from infinity at the object side into the treatment plane on the image side.
This means that the precision achievable in the therapy is determined by the precision accuracy on the one hand, but also by the size of the interaction zone on the other. The size of the interaction zone, in turn, with a given laser pulse width, is essentially defined by the size of the focus volume. The smaller the focus volume, the smaller is the interaction zone, and the lesser is the risk of damage to the surrounding tissue, because with a small focus volume, the photon density needed for the treatment effect is achieved already with a very low laser pulse energy, such as about 10 nJ to 200 nJ; as a result, the energy input to the in the vicinity of the interaction zone is low.
The size of the focus volume varies with the parameters of the optical system and with the wavelength of the therapeutic laser radiation. In other words: in connection with a given wavelength, the desired small size of the focus volume is made possible by small aberrations and a fairly high numerical aperture. With increasing numerical aperture, not only the lateral dimension of the focus shrinks, but also its axial dimension. From the viewpoint of application, the numerical aperture should preferably be as high as possible.
The possibilities known in prior art of medical treatment of the cornea lying at the periphery of the eye on the one hand, and of the crystalline lens lying within the eye on the other, have the disadvantage that the instruments available satisfy the requirements of their respective special purpose only, which means that they differ, especially with regard to the focus position in the eye, the aperture and the size of the focus volume, to such an extent that they are designed and suitably either for therapy of the cornea alone or for therapy of the crystalline lens alone.
This requires extensive instrumentation that is ineffective both with regard to purchase costs and because several separate instruments are used below their capacity most of the time. In addition, setting up the several instruments separately for examining and treating the same patient eye is time-consuming.
While laser therapy instruments that can treat both the crystalline lens and the cornea are known, they are originally optimized only for the treatment of the crystalline lens. They can be used, e.g., to make access cuts for cataract operations, but the precision achievable with them is insufficient for creating a flap. This is because axial focus movements by several millimeters are required if the laser focus is to reach the entire anterior segment of the eye including the crystalline lens.
An essential problem to be solved in that respect is the fact that all object-side movements serving to vary the focus position (varying the parameters of the therapeutic laser beam before it enters the optical focussing system, e.g., by the shifting of lenses within the optical systems arranged further up the beam or by the movement of scanning mirrors) will inevitably result in a change of beam paths within the optical focussing system. The term “optical focussing system”, in this context, stands for the objective from which the therapeutic laser radiation exits and is focussed on and directed at the eye. If, e.g., the axial position of the laser focus is shifted in this way, the aberrations occurring as a function of this variation will have a disadvantageous effect on the focus volume.
An optical focussing system or an objective that is optimized only for a particular focus position in axial direction always is a compromise between the spatial region accessible by the focus position and the size of the aberrations occurring within this region. As the same is true also for the lateral extension of the spatial region accessible by the laser focus, there is always a restriction of the entire spatial region in which the necessary focus quality is to be achieved.